Systems and methods for seeking sub-surface temperature conditions during tissue ablation

ABSTRACT

Systems and methods for ablating body tissue use an ablation element for contacting tissue to form a tissue interface. The ablation element is adapted to be connected to a source of ablation energy to conduct ablation energy for transmission by the ablation element into tissue at the tissue interface. The systems and methods include a tissue temperature sensing element held in a carrier in thermal conductive contact with tissue beneath the tissue interface. A mechanism attached to the carrier selectively advances the carrier relative to the ablation element to different depths beneath the tissue interface. A controller is coupled to the mechanism and to the tissue temperature sensing element to control advancement of the carrier beneath the tissue interface based, at least in part, upon tissue temperatures sensed by the sensing element beneath the tissue interface. Preferably, the controller controls the mechanism to locate the sensing element at the depth where the hottest sensed tissue temperature exists.

This is a continuation of application(s) Ser. No. 08/432,001 filed onMay 1, 1995, now abandoned.

FIELD OF THE INVENTION

In a general sense, the invention is directed to systems and methods forcreating lesions in the interior regions of the human body. In a moreparticular sense, the invention is directed to systems and methods forablating heart tissue for treating cardiac conditions.

BACKGROUND OF THE INVENTION

Physicians frequently make use of catheters today in medical proceduresto gain access into interior regions of the body. In some procedures,the catheter carries an energy transmitting element on its distal tip toablate body tissues.

In such procedures, the physician must establish stable and uniformcontact between the energy transmitting element and the tissue to beablated. Upon establishing contact, the physician must then carefullyapply ablating energy to the element for transmission to the tissue.

The need for precise control over the emission of ablating energy isespecially critical during catheter-based procedures for ablating hearttissue. These procedures, called electrophysiology therapy, are becomingincreasingly more widespread for treating cardiac rhythm disturbances,called arrhythmias. Cardiac ablation procedures typically use radiofrequency (RF) energy to form a lesion in heart tissue.

The principal objective of the invention is to provide systems andmethods for monitoring and reliably controlling the application ofenergy to ablate body tissue, thereby providing therapeutic results in aconsistent and predictable fashion.

SUMMARY OF THE INVENTION

The invention provides systems and methods that provide reliable controlover tissue heating and ablation procedures using temperature sensing.

The invention provides systems and methods for ablating body tissueusing an ablation element for contacting tissue to form a tissueinterface. The ablation element is adapted to be connected to a sourceof ablation energy to conduct ablation energy for transmission by theablation element into tissue at the tissue interface. The systems andmethods include a tissue temperature sensing element held in a carrierin thermal conductive contact with tissue beneath the tissue interface.

According to the invention, the systems and methods include a mechanismattached to the carrier to selectively advance the carrier relative tothe ablation element to different depths beneath the tissue interface. Acontroller is coupled to the mechanism and to the tissue temperaturesensing element to control advancement of the carrier beneath the tissueinterface based, at least in part, upon tissue temperatures sensed bythe sensing element beneath the tissue interface.

In a preferred embodiment, the controller controls the mechanism tolocate the sensing element at the depth where the hottest sensed tissuetemperature exists.

In a preferred embodiment, the ablation element is cooled.

In a preferred embodiment, the systems and methods include a secondcontroller coupled to the tissue temperature sensing element. The secondcontroller controls either the supply of ablation energy to the ablationelement, or the rate at which the ablation element is cooled, or both,based, at least in part, upon temperature sensed by the tissuetemperature sensing element.

In a preferred embodiment, the carrier holds the tissue temperaturesensing element in thermal conductive contact with tissue, while keepingthe temperature sensing element in isolation from the thermal conductivecontact with the electrode. The carrier has prescribed thermalconductive characteristics that significantly improve the sensitivity ofthe temperature sensing element to tissue temperature and not thetemperature of the ablation element.

Other features and advantages of the inventions are set forth in thefollowing Description and Drawings, as well as in the appended claims.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1A is a system for ablating tissue using an actively cooledablation electrode and associated cooling medium delivery system thatembodies the features of the invention;

FIG. 1B is a diagrammatic view of a lesion profile, without an activelycooled ablation electrode;

FIG. 1C is a diagrammatic view of a lesion profile, with an activelycooled ablation electrode;

FIG. 1D is a graph showing the increase in lesion volume as a functionof the cooling temperature of the ablation electrode;

FIG. 2A is a side section view of an actively cooled electrode of theopen system variety that can be used in the system shown in FIG. 1A;

FIG. 2B is a section view of the end of the actively cooled electrodeshown in FIG. 2A;

FIG. 3A is a side section view of another actively cooled electrode ofthe open system variety that can be used in the system shown in FIG. 1A;

FIG. 3B is a section view of the end of the actively cooled electrodeshown in FIG. 3A taken generally along line 3B--3B in FIG. 3A;

FIG. 4 is a diagrammatic view of an actively cooled electrode like thatshown in FIG. 3A in contact with tissue and with different types ofmedium being conveyed out of the cooling lumens;

FIG. 5 is a side section view of another actively cooled electrode ofthe open system variety that can be used in the system shown in FIG. 1A;

FIG. 6 is a side section view of an actively cooled electrode of theclosed system variety that can be used in the system shown in FIG. 1A;

FIG. 7 is a side section view of an electrode actively cooled using aPeltier diode that can be used in the system shown in FIG. 1A;

FIG. 8 is a diagrammatic view of a system for establishing a desiredtemperature boundary condition between an ablation electrode andendocardial tissue by actively cooling the electrode at a controlledrate;

FIG. 9A is a diagrammatic view of a system that adjusts the level of RFpower delivered to a cooled electrode based upon the sensed electrodetemperature and the rate that ablation power is conveyed into the tissuethrough the cooled electrode;

FIG. 9B is a diagrammatic view of a neural network that can be used inassociation with the system shown in FIG. 9A;

FIG. 10 is a side section view of an actively cooled energy transmittingelectrode that can be associated with the system shown in FIG. 1A,showing an outward projecting, blunt end temperature sensing elementcarried within a heat conducting cap by the electrode for sensing tissuetemperature below the tissue surface;

FIG. 11 is an exploded side view of the temperature sensing elementshown in FIG. 10;

FIG. 12 is a side section view of an actively cooled energy transmittingelectrode that can be associated with the system shown in FIG. 1A,showing an outward projecting, pointed end temperature sensing elementcarried within a heat conducting cap by the electrode for sensing tissuetemperature below the tissue surface;

FIG. 13 is a side section view of an actively cooled energy transmittingelectrode that can be associated with the system shown in FIG. 1A,showing a movable temperature sensing element carried within a heatconducting cap by the electrode, the sensing element being shown in itsretracted position within the electrode;

FIG. 14 is a side section view of the energy transmitting electrodeshown in FIG. 13, showing the movable temperature sensing element in itsextended position projecting into tissue;

FIG. 15A is a section view of the manually rotatable stylet used toadjust the position of the movable temperature sensing element shown inFIGS. 13 and 14;

FIG. 15B is an enlarged view of an actively cooled electrode with anexternally threaded, movable temperature sensing element;

FIG. 15C is an enlarged view of an actively cooled electrode with a corkscrew-type carrier for the temperature sensing element;

FIG. 16 is a section view of an alternative manual, push-pull typestylet used to adjust the position of the movable temperature sensingelement;

FIG. 17A is an enlarged end view of an actively cooled energytransmitting electrode carrying an outward projecting temperaturesensing element with multiple temperature sensors for sensing multiplesub-surface tissue temperatures;

FIG. 17B is an enlarged end view of an actively cooled electrodecarrying a temperature sensing element that projects into tissue havingmultiple temperature sensors associated with spaced regions of thermalconductive material substantially isolated from thermal conductivecontact with each other;

FIG. 18 is an enlarged end view of an actively cooled energytransmitting electrode carrying multiple temperature sensing elements,each sensing element projecting into tissue to sense sub-surface tissuetemperature;

FIG. 19 is a diagrammatic view of a system that adjusts the level of RFpower delivered to a cooled electrode based in part upon actual maximumsub-surface tissue temperatures sensed by a temperature sensing elementthat penetrates below the tissue surface;

FIG. 20 is a section view of a motor driven stylet used to adjust theposition of the movable temperature sensing element shown in FIGS. 13and 14, with an associated feedback controller that seeks the region ofhighest sub-surface tissue temperature;

FIG. 21 is a diagrammatic view of an apparatus for acquiringexperimental data to create a function that correlates a relationshipamong lesion boundary depth, ablation power level, ablation time,maximum tissue temperature, and electrode temperature that can be usedby a processing element to control an ablation procedure to targetlesion characteristics; and

FIG. 22 is a diagrammatic flow chart showing a process that the feedbackcontroller for the motor driven stylet shown in FIG. 20 can use toposition the temperature sensor in the region of highest sub-surfacetissue temperature.

The invention may be embodied in several forms without departing fromits spirit or essential characteristics. The scope of the invention isdefined in the appended claims, rather than in the specific descriptionpreceding them. All embodiments that fall within the meaning and rangeof equivalency of the claims are therefore intended to be embraced bythe claims.

DESCRIPTION OF THE PREFERRED EMBODIMENTS

FIG. 1A shows a system 10 for ablating human tissue that embodies thefeatures of the invention.

In the illustrated and preferred embodiment, the system 10 includes agenerator 12 that delivers radio frequency energy to ablate tissue. Ofcourse, other types of energy can be generated for tissue ablatingpurposes.

The system 10 also includes a steerable catheter 14 carrying a radiofrequency transmitting ablation electrode 16. In the illustratedembodiment, the ablation electrode 16 is made of platinum/iridium. Theablation electrode 16 can be made from other energy transmittingmaterials like, for example, stainless steel, gold, or silver.

In the illustrated embodiment, the system 10 operates in a unipolarmode. In this arrangement, the system 10 includes a patch electrode thatserves as an indifferent electrode 18. In use, the indifferent electrode18 attaches to the patient's back or other exterior skin area.

Alternatively, the system 10 can be operated in a bipolar mode. In thismode, the catheter 14 carries both electrodes.

The system 10 can be used in many different environments. Thisspecification describes the system 10 when used to provide cardiacablation therapy.

When used for this purpose, a physician steers the catheter 14 through amain vein or artery (typically the femoral vein or artery) into theinterior region of the heart that is to be treated. The physician thenfurther manipulates the catheter 14 to place the electrode 16 intocontact with the tissue within the heart that is targeted for ablation.The user directs radio frequency energy from the generator 12 into theelectrode 16 to ablate and form a lesion on the contacted tissue.

I. THE ABLATION CATHETER

In the embodiment shown in FIG. 1A, the catheter 14 includes a handle20, a flexible catheter body 22, and a catheter distal section 24, whichcarries the electrode 16.

The handle 20 encloses a steering mechanism 26 for the catheter tip 24.A cable 28 extending from the rear of the handle 20 has plugs 30. Someof the plugs 30 are coupled to a signal wire 32 (see FIG. 2A) thatextends from the ablation electrode 16 through the catheter body 22. Theplugs 30 connect to the generator 12 for conveying radio frequencyenergy to the ablation electrode 16 through the wire 32.

Left and right steering wires 34 (also see FIG. 2A) extend through thecatheter body 22 to interconnect the steering mechanism 26 in the handle20 to the left and right sides of a deflecting spring element 36.Rotating a steering lever 38 on the handle to the left causes thesteering mechanism 26 to pull on the left steering wire, causing thespring element 36 to bend to the left (as shown in phantom lines in FIG.1A). Similarly, rotating the steering lever 38 to the right causes thesteering mechanism 26 to pull on the right steering wire 34, causing thespring element 36 to bend to the right (as also shown in phantom linesin FIG. 1A). In this way, the physician steers the ablation electrode 16into contact with the tissue to be ablated.

Further details of this and other types of steering mechanisms for theablating element 10 are shown in Lunquist and Thompson U.S. Pat. No.5,254,088, which is incorporated into this Specification by reference.

A. Actively Cooled Electrodes

In the illustrated and preferred embodiment, the system 10 includes anassembly 40 for actively cooling the electrode 16. Cooling forces theelectrode-tissue interface to lower temperature values, As a result (asFIGS. 1B and 1C show), the hottest iso-thermal region T_(MAX) is shifteddeeper into the tissue. This, in turn, shifts the 50° C. iso-thermalregion (designated T_(50C)), which determines the boundary of the tissuerendered nonviable by ablation, deeper into the tissue. An electrodethat is actively cooled can be used to transmit more ablation energyinto the tissue, compared to the same electrode that is not activelycooled. As a comparison of FIGS. 1B and 1C shows, the net result isthat, with cooling, the lesion (designated L1 and L2, respectively, inFIGS. 1B and 1C) extends deeper into the tissue and has a larger volume.

FIG. 1D shows this effect graphically. Assuming a maximum tissuetemperature T_(MAX) of about 94° C., actively cooling the electrode toan electrode temperature T1 below about 35° C. leads to at least a 50%increase in lesion volume. At an electrode temperature T1 below about25° C., lesion volumes increase by about 100%, i.e., lesion volumesdouble in size.

There are various ways to structurally provide an electrode that can beactively cooled during use.

1. Open Loop Cooling

In the embodiment shown in FIGS. 2A and 2B, the catheter body 22includes an interior lumen 42. The proximal end of the lumencommunicates with a connection port 44 in the handle (see FIGS. 1A and15A). The distal end of the lumen 42 communicates with a hollow cavity46 formed in the electrode 16.

In the illustrated and preferred embodiment, the cavity 46 includes anarray of outlet apertures 48 clustered at the distal tip of theelectrode 16. Alternatively, a single centrally located outlet aperture,or other arrangements of one or more apertures, could be provided in thedistal tip of the electrode 16.

In this arrangement, the cooling assembly 40 includes a source 50 (seeFIG. 1A also) of a biocompatible medium, such as saline, with or withoutheparin. A mechanism 52 cools the medium source 50 to a desiredtemperature. A supply line 54 with an in-line pump 56 supplies thecooled medium to the connection port 44 on the handle 20. The cooledmedium flows through the lumen 42 and into the electrode cavity 46. Theoutlet apertures 48 discharge the cooled medium into the regionsurrounding the electrode, as FIG. 2A shows. Because the cooling mediumis discharged directly into the space surrounding the electrode 16, thisarrangement will be called "open" loop cooling.

The flow of cooling liquid through the electrode cavity 46 conveys heataway from the thermal mass of the electrode 16 by conductive andconvective cooling. The system further includes a controller 58 (seeFIG. 1A) for controlling the rate of cooling, as will be described ingreater detail later.

Preferably, the flow of media through the outlet apertures 48 issufficient to sustain a positive fluid pressure throughout use, therebypreventing clotting about the electrode 16. The size and number ofoutlet apertures 48 determine the magnitude of the flow resistancethrough the electrode 16.

The orientation of the outlet apertures 48 also affects the efficiencyof the cooling effect. Preferably, the outlet apertures 48 are clusteredat the distal end of the electrode 16, as FIGS. 2A and 2B show. Thisorientation directs the cooling medium through the entire length of theelectrode 16 for better cooling effect. The discharged cooling mediumalso flows directly into and through the electrode-tissue interface,causing direct cooling of the tissue area being ablated. The directcooling can reduce the incidence of charring.

FIGS. 3A and 3B show an alternative structural embodiment of an activelycooled electrode of an "open" loop type. In this embodiment, an exteriorsleeve 60 surrounds the catheter body 22, forming a circumferentialspace. The space is compartmentalized by dividers 62 (see FIG. 3B) intomultiple, circumferentially spaced lumens 64. Of course, the number oflumens 64 can vary.

The proximal end of the sleeve 60 communicates with the connection port44 on the handle 20. The lumens 64 simultaneously conduct cooling mediumsupplied to the connection port 44 by the source 50 via the supply line54 and in-line pump 56. The distal end of the sleeve 60 opens along theexterior sidewall of the electrode 16. There, the lumens 64 dischargethe cooling medium along the periphery of the electrode 16 to cool it.

Alternatively, the sleeve 60 can be made to be moved axially along thecatheter body 22 like an introducer sheath. In this arrangement, theposition of the slidable sleeve can be adjusted to achieve optimaloutflow of cooling medium about the electrode.

Optionally, as FIG. 4 shows, the multiple lumens 64 formed within theexterior sleeve 60 can conduct media having different characteristicsbenefiting the ablation process. For illustrative purposes, FIG. 4 showsthree lumens, designated 64A, 64B, and 64C. The lumen 64A adjacent theregion of the electrode 16 in most intimate contact with tissue conductsa hypertonic liquid A having a relatively low resistivity at, forexample, about 15 ohm·cm, compared to resistivity of blood, which isabout 150 ohm·cm. The hypertonic liquid A discharged in this regiontherefore improves the transmission of RF energy from the electrode 16into the tissue, regardless of whether or not the electrode 16 is alsoactually cooled by the liquid in the process. The other lumens 64B and64C adjacent the region of the electrode 16 exposed to the blood poolcan conduct another liquid B having a relatively high resistivity,compared to blood, of, for example, about 1500 ohm·cm. The liquid Bcould comprise, for example, a 5% dextrose solution. The liquid Btherefore reduces the transmission of RF energy from the electrode 16into the blood pool, again regardless of whether liquid B also cools theelectrode 16 in the process. Furthermore, heparin could be supplied withliquid A through the lumen 64A adjacent the tissue-contacting region ofthe electrode 16 to locally reduce the incidence of clotting, while noheparin is supplied through the lumens 64B and 64C adjacent theblood-pool exposed region of the electrode 16. In this way, the volumeof anticoagulant introduced into the blood pool can be more locallydirected and controlled.

FIG. 5 shows another alternative embodiment of an actively cooledelectrode of the "open" loop type. In this embodiment, the electrode 16comprises a foam body 66 of an open cell porous material coated with anelectrically conductive substance. The electrically conductive substancecan be coated on the porous material 66 using, for example conventionalion beam assisted deposition (IBAD), or similar vapor depositiontechniques.

Coated foam body 66 is molded to assume what can be called a normalshape. In the illustrated embodiment (as FIG. 5 shows), the normaluncompressed shape is generally spherical. However, the originaluncompressed shape can be rectangular, square, oval, toroid, orvirtually any other shape. Due to its porous, open structure, the body66 can be collapsed, without damage, by an external compression forceduring deployment in a guide tube (not shown) into another more compactshape.

As in the FIGS. 2A/B embodiment, an interior lumen 68 supplies coolingmedium to the porous material of the body 66 from an external source 50(not shown in FIG. 5). The porous material of the body 66 uniformlyperfuses the cooling medium from the lumen 68 for discharge at thesurface of the medium.

2. Closed Loop Cooling

FIG. 6 shows an embodiment of electrode 16 that is actively cooled in a"closed" loop manner. During "closed" loop cooling, the cooling mediumis not discharged outside the electrode 16 at the ablation site.Instead, the cooling medium is circulated back to the source 50 or towaste 70 away from the ablation site.

In this arrangement, the system includes, in addition to the previouslydescribed source 50, supply line 54, and pump 56, a return line 72 thatconveys medium away from the electrode 16. The catheter body 22 includesan interior supply lumen 74 and an interior discharge lumen 76. Theproximal ends of the lumens 74 and 76 communicate with the connectionport 44 on the handle 20, with the supply lumen 74 in communication withthe supply line 54 and the discharge lumen 76 in communication with thereturn line 72.

The distal ends of the lumens 74 and 76 communicate with a hollow cavity78 formed in the electrode 16. The supply line 54 supplies the cooledmedium through the supply lumen 74 into the cavity 78, while the returnline 72 returns the medium through the discharge lumen 76 to the mediumsource 50 or to waste 70. As before, the flow of cooling liquid throughthe electrode cavity 78 conveys heat away from the thermal mass of theelectrode by conductive and convective cooling.

In a "closed" loop arrangement, a pressurized gas could be used as thecooling medium. The pressurized gas would be allowed to expand withinthe electrode chamber, cooling the electrode by the Joule-Thompsoneffect. The use of a pressurized gas and the Joule-Thompson effect tocool an electrode is disclosed in Jackson et al. U.S. Pat. No.5,281,217, which is incorporated herein by reference.

3. Diode Cooling

In the alternative embodiment shown in FIG. 7, the cooling assembly 40includes a conventional Peltier cooling diode 80 associated with theelectrode 16, which is also electrically coupled by wire 32 to thegenerator 12. The materials of the diode 80 are complex alloys, onedoped "p" and the other doped "n", like semiconductors, creating athermocouple at the junction. An applied voltage potential passescurrent from a source 88 through the junction. The polarity of thevoltage creates a "cold" side 82 of the diode 80, which is coupled inthermal conductive contact to the electrode 16, and a "hot" side 84 ofthe diode 80, which is coupled in thermal conductive contact to a heatdispersing element 86. The dispersing element 86 can be carried on thecatheter body 22 away from the electrode 16 and in contact with theblood pool.

The passage of current through the diode 80 creates a heat pump fromcold side 82 to hot side 84, conducting heat energy from the thermalmass of the electrode 16 to the heat dispersing element 86. Heat energycan thus be transferred from the thermal mass of the electrode 16 tocool it.

FIG. 7 shows the Peltier diode 80 being used in place of a source 50 ofcooling medium to actively cool the electrode 16 in either an open orclosed loop fashion. It is believed that the heat transfer capabilitiesof conventional Peltier diodes 50, coupled with the normal convectivecooling effects of the dispersing element 86 by the blood pool canaccommodate the requirements for active cooling of most ablationelectrodes 16. Alternatively, the Peltier diode 80 can be used incombination with a flowing source 50 of cooling medium to actively coolthe electrode.

B. Ablation Control Using Electrode Cooling

1. Prescribing a Desired Lesion Depth

FIG. 8 diagrammatically shows a system 90 for establishing a desiredtemperature boundary condition between an ablation electrode 16 andendocardial tissue by actively cooling the electrode 16 at a controlledrate.

The system 90 includes the generator 12 of RF ablation energyelectrically coupled by the wire 32 to the ablation electrode 16, whichis deployed within the body in contact with heart tissue. In theillustrated embodiment, when used for cardiac ablation, the generator 12is typically conditioned to deliver up to 150 watts of power at a radiofrequency of 500 kHz.

The system 90 shown in FIG. 8 also includes the source 50 of medium forcooling the electrode, as well as the mechanism 52 for cooling themedium. The mechanism 52 includes a controller 92 for establishing andmaintaining a desired temperature for the cooling medium in the source.

The supply line 54 and in-line pump 56 provide communication between thesource 50 and the connection port 44 on the catheter handle 20.Operation of the pump 56 conveys the cooled medium to the electrode 16,as already described. FIG. 8 shows an open loop arrangement of the typeshown in FIGS. 3A/B. A controller 94 coupled to the pump 56 establishesand maintains a commanded flow rate. In a closed loop system, a returnline 72 conveys the medium from the electrode for return to the source50 or to waste 70, in the manner shown in FIG. 6.

As shown in FIG. 8, the electrode 16 carries a temperature sensor 96.The sensor 96 senses instantaneous temperatures (T1) of the thermal massof the electrode 16. The temperature T1 at any given time is a functionof the power supplied to the electrode 16 by the generator 12 and therate at which the electrode 16 is cooled by the medium.

The characteristic of a lesion can be expressed in terms of the depthbelow the tissue surface of the 50° C. isothermal region T_(50C), whichmarks the boundary of tissue rendered nonviable. FIG. 8 designates thisdepth as D_(50C). The depth D_(50C) is a function of the physicalcharacteristics of the ablation electrode (that is, its electrical andthermal conductivities and size); the angle between the tissue and theelectrode; the temperature T1 of the thermal mass of the electrode; themagnitude of RF power (P) transmitted by the electrode into the tissue,and the time (t) the tissue is exposed to the RF power. Theserelationships can be observed empirically and/or by computer modelingunder controlled real and simulated conditions, as the following Examplewill illustrate.

For a desired lesion depth D_(50C), additional considerations of safetyconstrain the selection of an optimal operating condition among theoperating conditions listed in the matrix. The principal safetyconstraints are the maximum tissue temperature T_(MAX) and maximum powerlevel P_(MAX).

The maximum temperature condition T_(MAX) lies within a range oftemperatures which are high enough to provide deep and wide lesions(typically between about 90° C. and 98° C.), but which are safely belowabout 100° C., at which tissue desiccation or tissue boiling is known tooccur. It is recognized that T_(MAX) will occur a distance below theelectrode-tissue interface between the interface and D₅₀.

The maximum power level P_(MAX) takes into account the physicalcharacteristics of the electrode and the power generation capacity ofthe RF generator 12.

EXAMPLE (Determining a D_(50C) Function)

A 3-D finite element model is created for a cooled 8F diameter/5 mm longablation electrode held generally perpendicular in contact with anapproximately 4 cm thick rectangular slice of cardiac tissue. The tip ofthe electrode extends about 1.3 mm into the tissue. The overall volumeis a parallelpiped 8 cm long, 4 cm wide, and 4 cm thick. The model has8144 nodes, using hexahedral elements and a nonuniform mesh.

The current density boundary conditions are set at the electrode, sothat after 120 seconds (t) the maximum tissue temperature (T_(MAX))reaches about 95° C. On the outer surface of the overall volume thepotential is set to zero, and the temperature is fixed at 37° C. toaccount for the average body temperature. At the nodes on the electrodesurface the temperature is set to a value that modeled the effects ofactively cooling the electrode tip. This value (T1) is varied between 4°C. and 50° C. The finite element convective boundary condition at theelectrode-blood interface is set to 1.8×10⁻⁵ Joule (J) per cubicmillimeter (mm³) second (s) Kelvin K (J/mm³ ·s·K).

COSMOS is used on a Hewlett Packard workstation to solve theelectrical-thermal equations. The analysis looks at the effects ofelectrode cooling on lesion volume, on radio frequency power (P)required to keep T_(MAX) at about 95° C., and on the distance of thehottest tissue region beneath the tissue-electrode interface. The lesiondimensions are estimated from the volume enclosed by the 50° C.isothermal surface.

The model results are corroborated with experimental data acquired usingthe apparatus shown in FIG. 21. A 4 cm thick slice of bovine heart H isfixed in good contact with a 144 cm² patch electrode EP inside a tank Tfilled with saline at 37° C. An ablation catheter C carrying a cooled 8Fdiameter/5 mm long electrode E is placed in contact with the tissuesurface H at an angle of 90°. Water at about 4° C. is circulated from asource CS inside the catheter. A 0.55 mm bead thermistor TM1 is placedat the electrode tip (to sense T1), and the sensed temperature (T1) isused as output to manually control the cooled water flow rate (as shownby dotted lines in FIG. 21). The sensed temperature (T1) is keptconstant at a predetermined value between 27° C. and 40° C. A secondthermistor TM2 is placed in the cardiac tissue H about 2 mm beneath theelectrode tip. The second thermistor TM2 placement corresponds to thehottest tissue temperature region predicted by the finite elementsimulations. The two thermistor readings are acquired at a sampling rateof 20 ms by LabView running on a Power Mac IIci. A 500 kHz sinusoidalsignal is applied between the ablation and indifferent electrodes usinga 150 W RF ablation system AS. The delivered RF power (P) is keptconstant at predetermined values between 6 watts (W) and 20 W.

After the experiments are completed, the heart is removed from the tank,sliced transversely at each of the lesions, and the dimensions of thecontours marking tissue discoloration are measured. The bovine tissueused typically discolors at about 60° C., so the values obtainedunderestimate the dimension of in vivo lesions consisting of tissueheated above 50° C.

The following matrix sets forth the D_(50C) function obtained using theabove described methodology.

                  TABLE 1                                                         ______________________________________                                        D.sub.50c Boundary Function                                                   t = 120 seconds and T.sub.MAX = 95° C.                                 (For 8 F 5 mm ablation electrode)                                                                Lesion     Distance                                                           Volume     to T.sub.MAX                                                                         P                                        T1 (° C.)                                                                        D.sub.50c (mm)                                                                         (mm.sup.3) (mm)   (W)                                      ______________________________________                                         4        10.2     1.25       1.51   37                                       10        10.1     1.19       1.4    35.4                                     20        9.8      1.13       1.24   33                                       25        9.7      1.04       1.18   32                                       30        9.2      0.99       1.08   31                                       37        9        0.89       0.97   29.5                                     50        8.8      0.9        0.78   26.5                                     ______________________________________                                    

Other matrices can be developed using the above-described methodologyfor an array of values for t and T_(MAX) to further define the D_(50C)function.

The function in Table 1 can be further supplemented by other empiricallyderived information showing the cooling media flow rate needed to obtaindifferent electrode temperatures for the particular electrode, as thefollowing Table 2 exemplifies:

                  TABLE 2                                                         ______________________________________                                        Average Flow Rate of Cooling Media (Cooled Water)                             vs. Electrode Temperature T1 at Constant Power                                Conditions                                                                    (For 8 F 5 mm ablation electrode)                                             T1 (° C.)                                                                       30           35        40                                            ______________________________________                                        Average  9.3 ml/min   5.3 ml/min                                                                              4 ml/min                                      Flow                                                                          ______________________________________                                    

The system 90 includes a master controller 98. The master controller 98is coupled to the RF generator 12, the temperature sensor 96, thecooling controller 92, and the pump controller 94. The master controller98 includes in memory a matrix of operating conditions defining theD_(50C) temperature boundary function, as described above for t=120seconds and T_(MAX) =95° C. and for an array of other operatingconditions.

The master controller 98 includes an input device 100. In the system 90shown in FIG. 8, the physician uses the controller input device 100 toset a desired lesion depth in terms of D_(50C). The physician also usesthe input device 100 to identify the characteristics of the electrode,using a prescribed identification code; set a desired maximum RF powerlevel P_(MAX) ; a desired time t; and a desired maximum tissuetemperature T_(MAX).

For example, assume that the physician selects an 8F/5 mm ablationelectrode. The physician also selects a desired therapeutic result interms of a lesion depth D_(50C) =9.2 mm. The physician further selectsother details of the desired therapeutic result in terms of a targetedablation time of t=120 seconds; a maximum tissue temperature T_(MAX)=95° C.; and a maximum ablation power level P_(MAX) =50 W.

Based upon these inputs, the master controller 98 compares the desiredtherapeutic result to the function defined in the matrix (as exemplifiedby the above Tables 1 and 2). The master controller 58 selects anoperating condition to achieve the desired therapeutic result withoutexceeding the prescribed T_(MAX) by controlling the function variables.

In this example, based upon a desired T_(MAX) of 95° C. and t=120seconds, the controller 98 commands the generator 12 to maintain a fixedpower level P of 31 W (which does not exceed P_(MAX)) for the prescribedtime t=120 seconds. The controller 98 simultaneously controls the rateat which the electrode 16 is cooled (based upon Table 2) to establishand maintain T1 at the level called for by the function for the D_(50C)=9.2 mm boundary selected, which in this example is T1=30° C. (flowrate=9.3 ml/min).

The maximum tissue temperature will continuously increase toward T_(MAX)during the targeted ablation period t, with the rate of increasedepending principally upon the magnitude of P and T1. That is, the rateof tissue temperature increase with be greater at higher values of P andlower values of T1, and vice versa.

The master controller 98 can control the cooling rate in various ways.For example, the master controller 98 can control the rate of cooling bycommanding the temperature controller 92 to adjust the temperature ofthe cooling medium over time in response to variations in T1 toestablish and maintain the set T1. Alternatively, the master controller98 can control the rate of cooling by commanding the pump controller 94to adjust the flow rate of the cooling medium over time in response tovariations of T1 to establish and maintain the set T1. The mastercontroller 98 can also command the controllers 92 and 94 in tandem toreach the same result.

The manner in which the master controller 98 processes informationpertaining to T1 to derive control signals to vary medium temperatureand medium flow rate can vary. For example, the master controller 98 canemploy proportional control principles, proportional integral derivative(PID) control principles, adaptive control, neural network, and fuzzylogic control principles.

When cooling is accomplished using the Peltier cooling diode 80 (as FIG.7 shows), the master controller 98 establishes and maintains T1 bycommanding the current source 88 to adjust current flow to the diode 80.When the diode 80 is used in combination with active medium flow cooling(as FIG. 8 shows), the master controller 98 can set the mediumtemperature and the medium flow rate in the manners above described, andfurther control the current source 88 to the diode to accomplish fineadjustments to maintain the desired T1.

It can be appreciated that various combinations of cooling control usingthe diode 80 are also possible. As before stated, the master controller98 can employ proportional control principles, proportional integralderivative (PID) control principles, adaptive control, neural network,and fuzzy logic control principles in varying the current flow to thediode 80 based upon changes of T1 over time.

At the end of the targeted ablation period t, the controller 98terminates power to the ablation electrode. The desired lesion depthwill be formed, and T_(MAX) will not have exceeded the target of 95° C.

In alternative arrangements, the controller 98 can fix any one or moreof the control variables T1, P, or t and vary the remaining one or moreof the control variables T1, P, or t to achieve the desired D_(50C)temperature boundary. The system 90 thereby permits the physician, ineffect, to "dial-a-lesion" by specifying a desired D_(50C). Using activecooling in association with time and power control, the controller 98achieves the desired D_(50C) without the need to sense actual tissuetemperature conditions.

2. Predicting Maximum Tissue Temperature/Depth During Cooling

FIG. 9A shows a system 102 that adjusts the level of RF power deliveredto a cooled electrode 16 and/or the cooling rate based upon a predictionof instantaneous maximum tissue temperature, which is designated Ψ_(MAX)(t).

In a preferred implementation, the prediction of Ψ_(MAX) is derived by aneural network, which samples at the current time (t) a prescribednumber (k_(n)) of previous power levels P, previous rates at which heathas been removed to cool the electrode, and previous electrodetemperature.

The heat removal rate is identified by the expression Å, where

    Å=c×ΔT×RATE

where:

c is the heat capacity of the cooling medium used (in Joules (J) perkilogram (kg) Kelvin (K), or J/kg K) ΔT is the temperature drop in thecooling medium during passing through the electrode 16 (K), and

RATE is the mass flow rate of the cooling medium through the electrode(kg/sec).

The heat transmitted by the ablation electrode to the tissue is thedifference between the heat generated by Joule effect and the heatremoved by active cooling. At a given temperature T1 and flow rate ofcooling medium, the magnitude of Å increases as RF power delivered tothe electrode 16 increases. Together, T1 and Å represent an indirectmeasurement of how rapidly the sub-surface tissue temperature ischanging. Together, T1 and Å are therefore predictive of the depth andmagnitude of the hottest sub-surface tissue temperature Ψ_(MAX), andthus indirectly predictive of the lesion boundary depth D_(50C). Largedeep lesions are predicted when T1 is maintained at a low relativetemperature (by controlling cooling rate) and Å is maintained at a highrelative value (by controlling RF power). Likewise, smaller lesions arepredicted when T1 is maintained at a high relative temperature and Å ismaintained at a low relative value.

The system 102 shown in FIG. 9A implements these control criteria usingan electrode 16 of a closed system type, like that shown in FIG. 6. Theelectrode 16 carries three temperature sensing elements 104, 106, and108. The first sensing element 104 is in thermal contact with thethermal mass of the electrode 16 to sense its temperature, or T1 asalready described. The second sensing element 106 is located to sensethe temperature of the cooling medium as it enters the electrode cavity78, or T_(IN). The third sensing element 108 is located to sense thetemperature of the cooling medium as it exits the electrode cavity 78,or T_(OUT). In this closed system arrangement, the temperature increasein the cooling medium during its passage through the electrode ΔT iscomputed as follows:

    ΔT=T.sub.OUT -T.sub.IN Closed System

In an open system arrangement (like that shown in FIGS. 2A/B and 3A/B),where the cooling medium is discharged directly in the region of tissuein contact with the electrode 16, there is no third temperature sensingelement 108. In this case, ΔT is computed as follows:

    ΔT=T1-T.sub.IN Open System

In systems where environmental variables are closely controlled, theprediction of Ψ_(MAX) may be derived from sampling at the current time(t) a prescribed number (k_(n)) of previous power levels P and previouselectrode temperatures, without sampling Å.

In FIG. 9A, the master controller 98 is coupled to the RF generator, thetemperature sensing elements 104, 106, and 108 (or 104 and 106 in anopen system), the cooling controller 92, and the pump controller 94.

The controller 98 includes a neural network predictor 144 (see FIG. 9B).The predictor 144 can comprise a two-layer neural network, although morehidden layers could be used. The predictor 144 receives as inputs afirst set of k₁ of weighted past samples of Å, {Å(t-1) to (t-k₁)}; asecond set of k₂ of weighted past samples of P, {P(t-1) to (t-k₂)}; anda third set of k₃ samples of T1, {T1(t-1) to (t-k₃)}. The number ofsamples in the sets k₁,2,3 can be varied, according to the degree ofaccuracy required. As an example, k₁ and k₂ are preferably in the rangeof 5 to 20. k₃ can be selected equal to 1.

The predictor 144 can be variously configured. In the illustratedembodiment, the predictor 144 comprises a two layer neural network,although more hidden layers could be used.

In this implementation, the predictor 144 includes first and secondhidden layers and four neurons, designated N.sub.(L,X), where Lidentifies the layer 1 or 2 and X identifies a neuron on that layer. Thefirst layer (L=1) has three neurons (X=1 to 3), as follows N.sub.(1,1) ;N.sub.(1,2) ; and N.sub.(1,3). The second layer (L=2) comprising oneoutput neuron (X=1), designated N.sub.(2,1).

The weighted past samples {Å(t-1) to (t-k₁)}; {P(t-1) to (t-k₂)}; and(in the alternative embodiment) T1, {T1(t-1) to (t-k₃)} are fed asinputs to each neuron N.sub.(1,1) ; N.sub.(1,2); and N.sub.(1,3) of thefirst layer.

The output neuron N.sub.(2,1) of the second layer receives as inputs theweighted outputs of the neurons N.sub.(1,1) ; N.sub.(1,2) ; andN.sub.(1,3). Based upon these weighted inputs, the output neuronN.sub.(2,1) outputs Ψ_(MAX) (t).

The predictor 144 must be trained on a known set of data that have beenpreviously acquired experimentally. For example, using aback-propagation model, the predictor 144 can be trained to predict theknown hottest temperature of the data set with the least error. Once thetraining phase is completed, the predictor 144 can be used to predictΨ_(MAX) (t).

Alternatively, fuzzy logic or linear prediction algorithms can be usedto derive Ψ_(MAX) (t) from sampling past power P, electrode temperatureT1, and (in the preferred embodiment) cooling rate Å.

The master controller 98 receives from the physician, via the inputdevice 100, a desired maximum tissue temperature value TT_(SET), adesired electrode temperature T1_(SET), and a P_(MAX).

The set temperature value TT_(SET) represents the desired hottestsub-surface tissue temperature that the physician wants to maintain atthe ablation site, consistent with the need to prevent micro-explosions.The value TT_(SET) can comprise a fixed, targeted magnitude, or thevalue of TT_(SET) can vary over time to define a set temperature curve,which can be either linear or nonlinear. Further details of using settemperature curves are disclosed in U.S. patent application Ser. No.08/266,023, filed Jun. 27, 1994, and entitled "Tissue Heating andAblation Systems and Methods Using Time-Variable Set Point TemperatureCurves for Monitoring and Control."

For T1_(SET), the preferred embodiment takes into account therelationship between electrode temperature T1 and increases in lesionvolume shown in FIG. 1D, selecting as the desired T1_(SET) a temperaturebelow about 25° C. and, most preferable, between about 10° C., and about25 C.

The value P_(MAX) is the highest allowed power level, based uponconsiderations already stated.

The master controller 98 periodically derives Ψ_(MAX) (t) and comparesΨ_(MAX) (t) to TT_(SET) (t). Based upon this comparison, the mastercontroller 98 derives a demand power output, taking into accountP_(MAX), while cooling to maintain T1_(SET). The demand power outputrepresents the magnitude of the radio frequency power that should besupplied to the electrode 16 to establish and contain the desiredmaximum tissue temperature TT_(SET) at a fixed value or along a setlinear or nonlinear curve.

Alternatively, the master controller 98 could maintain a fixed powerlevel below P_(MAX) and adjust the cooling rate Å based upon Ψ_(MAX) (t)to contain TT_(SET) at a fixed value or along a set curve. As beforedescribed, the master controller 98 can control the cooling rate bycommanding the temperature controller 92 to adjust the temperature ofthe cooling medium over time, or by commanding the pump controller 94 toadjust the flow rate of the cooling medium over time, or by commandingthe controllers 92 and 94 in tandem to reach the same result.

The manner in which the controller 98 derives the control commands canvary. For example, it can employ proportional control principles,proportional integral derivative (PID) control principles, adaptivecontrol, neural network, and fuzzy logic control principles. Furtherdetails of these control principle are disclosed in copending U.S.patent application Ser. No. 08/266,023, filed Jun. 27, 1994, andentitled "Tissue Heating and Ablation Systems and Methods UsingTime-Variable Set Point Temperature Curves for Monitoring and Control."

Using active cooling in association with power control and/or rate ofenergy removal at the electrode, the controller 98 achieves the desiredrate of energy removal Å to achieve a desired lesion characteristic.Like the system 90 shown in FIG. 8, the system 102 shown in FIG. 9Aachieves its lesion formation objectives without the need to senseactual tissue temperature conditions.

Alternatively, the master controller 98 can use a matrix function tocorrelate observed operating conditions, which Tables 1 and 2 exemplifyin partial form, to infer Ψ_(MAX) without actually sensing tissuetemperature conditions.

In this implementation, the controller 98 senses the flow rate ofcooling media, the sensed electrode temperature T1, and the power P. Thecontroller 98 compares these sensed values to values set forth by inmatrix function. The controller 98 infers from this comparison whatT_(MAX) would be, according to the function, under these sensedoperating conditions. The T_(MAX) inferred under this methodologybecomes Ψ_(MAX). For example, at a sensed cooling flow rate of 9.3ml/min, a sensed power P of 31 W, and a sensed electrode temperature T1of 30° C., Tables 1 and 2 would infer that T_(MAX) would be 95° C. at anablation time (t) of 120 seconds. In this implementation the inferredmaximum tissue temperature becomes Ψ_(MAX) ·Power and/or cooling rateare then controlled to contain Ψ_(MAX) at a fixed value or along a setcurve.

3. Sensing Actual Maximum Tissue Temperature/Depth During Cooling

In the embodiments shown in FIGS. 10 to 12, the cooled ablationelectrode 16 carries at least one temperature sensing element 110 forsensing actual tissue temperature. In these embodiments, the power thatthe RF generator 12 applies to the electrode 16 is set, at least inpart, by the actual tissue temperature conditions sensed by the element110.

In the illustrated embodiment, the temperature sensing element 110comprises a conventional small bead thermistor 112 with associated leadwires 114. In a preferred implementation, the thermistor 42 comprises a0.55 mm bead thermistor commercially available from Thermometrics(Edison, N.J.), Part Number AB6B2-GC16KA143E/37° C-A.

It should be appreciated that other types of temperature sensingelements can also be used. For example, a thermocouple could be used asthe temperature sensing element. In a preferred implementation, thethermocouples are constructed by either spot welding or by laserstripping and welding the different metals together to form thethermocouple junction. When a thermocouple serves as the temperaturesensing element, a reference thermocouple must be used. The referencethermocouple may be placed in the handle 20 or exposed to the blood poolin the manner disclosed in copending U.S. patent application Ser. No.08/286,937, filed Aug. 8, 1994, and entitled "Systems and Methods forSensing Temperature Within the Body."

Potting compound 116 encapsulates the thermistor 112 and lead wires 114.The lead wires 114 are also enclosed in insulating sheaths 117, whichelectrically isolate the wires 114. Together, the compound 116 andsheaths 117 electrically insulate the thermistor 112 from thesurrounding ablation electrode 16. For better performance, the wiresshould be electrically shielded.

The potting compound 116 and insulation sheaths 117 can be made withvarious materials. In the illustrated embodiment, heavy isomid serves asthe potting compound 116, although another cyanoacrylate adhesive, asilicon rubber RTV adhesive, polyurethane, epoxy, or the like could beused. The sheaths 117 are made from polyimide material, although otherconventional electrical insulating materials also can be used.

Similar electrical insulation is required when thermocouples are used asthe temperature sensors. For example, the thermocouple junction can beplaced in a thermally conducting epoxy inside a polyester sleeve. In apreferred implementation, the thermocouple junction is placed in siliconrubber RTV adhesive (NuSil Technologies, Carpenteria, Calif.) within ashrink polyester sleeve, which is then shrunk to fit tightly about thethermocouple junction and wires. To reduce electrical interference, thethermocouple wires are also preferably shielded and twisted together.

The lead wires 114 for the thermistor 112 extend through the catheterbody 22 and into the catheter handle 20 (see FIG. 15A). There, the leadwires 114 electrically couple to the cable 28 extending from the handle20. The cable 28 connects to the generator 12 and transmits thetemperature signals from the thermistor 112 to the generator 12.

In the embodiment illustrated in FIGS. 10 to 12, the ablation electrode16 includes an interior well 118 at its tip end. The temperature sensingelement 110 occupies this well 118. The sensing element 110 shown inFIGS. 10 to 12 extends beyond the tip of the electrode 16 to projectbeneath the surface of the endocardium. The sensing element 110 isthereby positioned to sense actual sub-surface tissue temperatureconditions.

In the illustrated and preferred embodiment, the sub-surface temperaturesensing element 110 is enclosed within a thermally conducting cap 120(see FIGS. 10 and 11). The cap 120 comprises a material having a highthermal conductivity that is at least 1.0 watt (W) per meter (m) Kelvin(K), or 1.0 W/m K. Metallic materials like stainless steel, gold, silveralloy, platinum, copper, nickel, titanium, aluminum, and compositionscontaining stainless steel, gold, silver, platinum, copper, nickel,titanium, and aluminum possess this degree of thermal conductivity. Forexample, stainless steel has a thermal conductivity of about 15 W/m K,and platinum has a thermal conductivity of about 71 W/m K. This thermalconductivity is significantly higher than the thermal conductivity ofconventional polymer potting material surrounding the temperature sensor110. For example, silicon rubber has a thermal conductivity of onlyabout 0.13 W/m K, and polyurethane has a thermal conductivity of onlyabout 0.026 W/m K.

The cap 120 has an open interior 122. The encapsulated thermistor 112snugly occupies the open cap interior 122 in thermal conductive contactwith the thermal conducting material of the cap 120. Preferably, thethermistor 112 is potted within the open interior 122 using an epoxyhaving an enhanced thermal conductivity that is at least 1.0 W/m K. Theinclusion of a metallic paste (for example, containing aluminum oxide)in a standard epoxy material will provide this enhanced thermalconductivity. When the ablation energy is radio frequency energy, thepotting material must also electrically insulate the temperature sensingelement 112 from the cap 120.

The cap 120 in turn is fitted within the well 118 of the electrode 16.The cap 120 has a distal end 124 that makes thermal conductive contactwith the tissue. The high thermal conductivity of the cap materialassures that the cap 120 will quickly reach an equilibrium temperatureclose to that of the tissue it contacts.

In a representative preferred implementation (see FIG. 3), the cap 120is made from stainless steel 304 (having a thermal conductivity of about15 W/m K). The cap 120 has a wall thickness along the sidewall and atthe distal end of about 0.005 inch. The cap 120 has an overall length ofabout 0.060 inch and an overall width of about 0.033 inch (the openinterior being about 0.022 inch in width). The encapsulated thermistor42 is fixed to the cap interior 56 using a thermally conducting epoxylike EP42HTAO (Master Bond, Inc., Hackensack, N.J.). The thermalconductivity of this epoxy (with aluminum oxide) is about 1.15 W/(m K).

The cap 120 provides enhanced thermal conducting characteristics,creating an isothermal surface around the sub-surface sensing element110 in thermal equilibrium with the surrounding tissue temperatureconditions. The cap 120 also provides added strength to resist bendingor fracture during manufacturing and handling.

In the illustrated and preferred embodiment, a thermal and electricallyinsulating barrier 142 forms an interface between the interior wall ofthe well 118 and the side of the cap 120 that occupies it. In apreferred embodiment, the barrier 142 comprises polyamide adhered aboutthe sidewall of the cap 120 using FMD-14 to serve as an electricalinsulator. The barrier 142 also comprises polyester shrink tubingsecured by heat shrinking about the polyamide to serve as a thermalinsulator.

In the illustrated and preferred embodiment, a thermal insulating tube144 also lines the interior of the well 118. The tube 144 furtherthermally insulates the temperature sensing element 40 from the thermalmass of the electrode 16. In the illustrated and preferred embodiment,the thermistor-containing cap 120 and associated barrier 142 are affixedby potting within the electrode well using cyanoacrylate FMD-13 (LoctiteCorporation, Newington, Conn.).

Therefore, the temperature condition sensed by the sensing element 40within the cap 120 closely represents the actual tissue temperaturecondition that the cap 120 contacts.

EXAMPLE

The thermal sensitivity of a temperature sensing element enclosed in athermally conductive carrier according to the invention (Sensor 1) wascompared to the thermal sensitivity of a temperature sensing elementfree of the carrier (Sensor 2).

Sensor 1 was carried within the well of an 8F 4 mm standardplatinum/iridium radio frequency transmitting electrode. Sensor 1comprised a 0.55 mm bead thermistor embedded in a glass bead, which inturn was embedded in an epoxy resin, which was encapsulated in apolyimide sheath. The entire encapsulated thermistor assembly wasmounted by FMD-14 within a cap, as above described, made of stainlesssteel 304 having a wall thickness of 0.005 inch. The exterior side wallsof the cap were thermally isolated from the electrode by one layer ofpolyamide and one layer of polyester shrink tubing. The assembly waspotted within the electrode well using FMD-13. The distal tip of the capwas free of thermal insulating material and was flush with the distaltip of the electrode for contact with tissue.

Sensor 2 comprised a thermocouple potted with solder in thermalconductive contact within an 8F/4 mm standard platinum/iridium radiofrequency transmitting electrode.

The thermal sensitivity of each Sensor 1 and 2 was tested by placing theconsolidated electrode and sensor assembly into a water bath maintainedat 20° C. A soldering wand maintained at a temperature of 60° C. wasplaced into contact with each electrode beneath the surface of thewater. This contact was maintained to achieve steady state conditionsboth against the side of the electrode (the electrode being heldhorizontally) and at the distal tip of the electrode (the electrodebeing held vertically). The temperatures sensed by each Sensors 1 and 2in both electrode orientations were recorded.

The following Table 3 summarizes the results:

TABLE 3

Comparison of the Thermal Sensitivity of a Temperature Sensor CarriedWithin a Thermal Conductive Surface to the Thermal Sensitivity of aTemperature Sensor Without a Thermal Conductive Surface

    ______________________________________                                                     VERTICAL                                                                              HORIZONTAL                                                            POSITION                                                                              POSITION                                                 ______________________________________                                        SENSOR 1 (With 59° C.                                                                           40° C.                                        Thermal                                                                       Conductive                                                                    Surf ace)                                                                     SENSOR 2       40° C.                                                                           39° C.                                        (Without Thermal                                                              Conductive                                                                    Surface)                                                                      ______________________________________                                    

The above Table shows that Sensor 2 is not sensitive to the actualtemperature of the 60° C. heat source. Regardless of its orientation,Sensor 2 continues to sense the 40° C. temperature of the thermal massof the electrode itself (the remainder of the heat energy of the sourcebeing dissipated by the surrounding water bath).

In contrast, Sensor 1 shows significant sensitivity with respect to itscontact orientation with the 60° C. heat source. When held horizontally,out of direct contact with the heat source, Sensor 2, like Sensor 1,senses the 40° C. temperature of the thermal mass of the electrodeitself. However, when held vertically, in direct contact with the heatsource, Sensor 1 essentially senses the actual temperature of the heatsource, and not the temperature of the electrode. The cap encapsulatingSensor 1, having a high intrinsic thermal conductivity of at least 1.0W/m K, directly conducts heat from the source for sensing by Sensor 1.The thermal conducting cap creates an isothermal condition about Sensor1 close to the actual temperature of the source. Furthermore, the cap,being substantially isolated from thermal conductive contact with theelectrode, retains this isothermal condition about Sensor 1, preventingits dissipation by the thermal mass of the electrode.

In quantitative terms, the 59° C. temperature sensed by Sensor 1 when indirect contact with the 60° C. heat source, compared to the 40° C.electrode temperature sensed when not in direct contact with the source,accounts for 19 of the total 20 units of actual temperature differencebetween the heat source and the electrode. Thus, in quantitative terms,the presence of the thermal conducting cap in Sensor 1 establishes a 95%sensitivity to the temperature of the heat source (i.e., which, in use,would be sensitivity to actual tissue temperature), and only a 5%sensitivity to the temperature of the electrode itself. This is comparedto an essentially 100% sensitivity of Sensor 2 to the temperature of theelectrode. In the absence of the cap that embodies the invention, Sensor2 is virtually insensitive to the actual temperature of the heat source(i.e., actual tissue temperature).

In the embodiment shown in FIG. 10, the cap 120 presents a blunt distalend 124 that projects from the end of the electrode 16, without actuallypenetrating it. As FIG. 10 shows, the endocardium is malleable enough toconform about the electrode 16 and the projecting cap 120.

In the alternative embodiment shown in FIG. 12, the cap 120 presents asharpened distal end 124 that actually penetrates the endocardium. Bycausing the cap 120 to actual penetrate the endocardium, better uniformtissue contact is achieved, both beneath the surface about thetemperature sensing element 110 and at the surface along the electrode.

The temperature sensing element 110 can project into the tissue at anydepth desired, depending upon the tissue morphology of the individualpatient and the experience and judgment of the attending physician,provided, of course, that transmural penetration of the heart wall doesnot occur.

In the preferred embodiment (see FIGS. 13 and 14), the temperaturesensing element 110 is movable by the physician between a retractedposition within the electrode well 118 (shown in FIG. 13) and anextended position outside the electrode well 118 (shown in FIG. 14)projecting into tissue. In FIGS. 13 and 14, the temperature sensingelement 110 is shown to have a blunt distal end 124, although a sensingelement 110 having a sharpened distal end could also be used.

The movable nature of the temperature sensing element 110 shown in FIGS.13 and 14 provides added protection against bending or fracture of theelement until the moment of use. The element 110 can be retained in aretracted, not exposed position during handling outside the body andwhile being deployed to the desired site within the body.

The movement of the temperature sensing element can be accomplished invarious ways. In the embodiment shown in FIGS. 13 and 14, a stylet 126extends through the catheter body 22 within a braided protective sleeve128 made of, for example, polyimide or stainless steel. The proximal endof the stylet 126 is attached to a control knob 130 on the handle 20(see FIG. 15A). The distal end of the stylet 126 is secured by adhesive,solder, crimping, or the like to the cap 120.

The thermistor wires 114 extend along the outside of the stylet 126within the protective sleeve 128 (see FIGS. 13 and 14). Another sleeve132 of electrically insulating material, like heat shrink tubing madefrom Teflon® or polyester material, surrounds the stylet 126 and wires114 up to and around the junction between the cap 120 and the stylet126. The sleeve 132 holds the wires 114 tightly against the stylet 126.The sleeve 132 also creates a smooth transition between the stylet 126and cap 120, while further provides protection against electricalinterference. A sleeve 136 of thermally insulating material, likepolyimide, also preferably lines the interior of the well, to thermallyinsulate the cap 120 from the thermal mass of the electrode 16.

The stylet 126 can be manually or automatically advanced in variousways. In the illustrated embodiment, the stylet 126 includes helicallands 138 formed along its length (see FIG. 15A). The lands 138 engagemating screw threads 142 within a stationary guide element 140 withinthe handle 20. Rotation of the control knob 130 by the physician rotatesthe stylet 126 within the guide element 140. Upon rotation in onedirection, the helical lands 142 advance the stylet forward axiallywithin the catheter body 22. Upon rotation in the opposite direction,the helical lands 142 move the stylet rearward axially within thecatheter body 22. In this way, the sensing element 110 can beincrementally moved in a controlled fashion between the retracted andextended positions.

In this arrangement (see FIG. 15B), the distal cap end 124 can itself bethreaded with helical lands 146. Upon rotational advancement of thesensing element 110 by the stylet 126, the helical lands 146 engagetissue to better anchor the element 110 for temperature sensing.Alternatively (see FIG. 15C), the stylet 126 can be attached to acarrier 150 configured as a cork-screw. Like the helical lands 146, thecork-screw carrier 150 engages tissue during rotation as the stylet 126is advanced forward by rotation. As FIG. 15C shows, the temperaturesensing element 110 is secured in thermal conductive contact with thecork-screw carrier 150 near its distal tip.

In the illustrated and preferred embodiment, the distal cap end 124 andthe distal tip of the electrode 16 are marked with a fluoroscopicallydense material. In this way, the travel of the temperature sensingelement 110 into the tissue can be monitored by fluoroscopy as thephysician incrementally advances the element 110.

Alternatively, the stylet 126 can be advanced without rotation. In thisarrangement (see FIG. 16), the proximal end of the stylet 126 includes aseries of ribs 152, which successively make releasable, snap-fitengagement with detent 154 in the handle 20. As the physician moves thestylet 126 in a linear (push-pull) direction, the detent 154 capturesthe ribs 152 one at a time, releasing the captured rib 152 in responseto further linear force. Like the rotating stylet 126 shown in FIG. 8,the linear (push-pull) stylet 126 shown in FIG. 16 permits controlled,incremental movement of the sensing element 110 into and out of tissuecontact.

In FIGS. 10 to 16, the actively cooled electrodes 16 shown are of themetal types shown in FIGS. 2A/B and 3A/B. It should be appreciated thata porous, actively cooled electrode body 66 like that shown in FIG. 5can also carry a temperature sensing element 110 of a fixed or movablekind.

In another alternative embodiment shown in FIG. 17A, the actively cooledelectrode 16 (which is of an open system type, having outlet apertures48 for the cooling medium like that shown in FIGS. 2A/B) includes atemperature sensing element 110 having multiple thermocouples designated112(1), 112(2), and 112(3). The multiple thermocouples 112(1), 112(2),and 112(3) are arranged in a housing 156 in a spaced-apart stackedrelationship along the axis of the housing 156. The housing 156 can befixed in an outwardly projecting position, as FIGS. 10 and 12, or thehousing 90 can be moved into an out of the projecting position in themanner of the stylet-movable cap 120 previously described (as shown inFIGS. 13 and 14).

In one embodiment (as FIG. 17A shows), the housing 156 comprises a bodyformed from a conventional potting compound, like silicon rubber, RTVadhesive, polyurethane, or epoxy, having a thermal conductivity lessthan the tissue it contacts. In the illustrated environment, where thethermal conductivity of myocardium is about 0.43 W/m K, pottingcompounds like silicon rubber and polyurethane material, for example,have thermal conductivities of, respectively, 0.13 W/m K and 0.026 W/mK. The relatively low thermal capacity of this material conditions theelements 112(1)/112(2)/112(3) to sense localized relative changes in thetissue temperature gradient along the length of the housing 156. Thesensing of the relative temperature gradient permits the identificationalong the gradient of the maximum tissue temperature region for controlpurposes, although the temperatures sensed by the elements112(1)/112(2)/112(3) will not directly represent actual tissuetemperatures.

If a more direct correspondence between sensed and actual tissuetemperatures is required, the housing 156 (see FIG. 17B) can includespaced bands 158(1), 158(2), and 158(3) of thermal conductive materialhaving thermal conductivity well above the contacted tissue, of at least1.0 W/m K, as already described. The spaced bands 158(1), 158(2), 158(3)establish localized regions of thermal conductive contact betweenindividual sensing element 112(1), 112(2), and 112(3) and tissueimmediately adjacent to the respective band. Thermal insulating material160 substantially isolates the spaced bands 112(1), 112(2), and 112(3)from thermal conductive contact with each another. The thermallyisolated bands 112(1), 112(2), and 112(3), each with a relatively highthermal conductivity, more accurately obtain the actual tissuetemperature gradient along the length of the housing 156, than whenmaterials with lower thermal conductivities are used.

In either embodiment, the multiple, axially stacked thermocouples112(1), 112(2), and 112(3) allow the physician to obtain and monitor aprofile of temperate conditions at different depths beneath the tissuesurface. The physician can manually select for ablation control purposesthe one thermocouple located in the hottest sub-surface temperatureregion. Alternatively, an automated control mechanism can automaticallycompare temperatures from all thermocouples 112(1), 112(2), and 112(3)and output the hottest sub-surface temperature for temperature controlpurposes.

In the embodiment shown in FIG. 18, an array of multiple, spaced-aparttemperature sensing elements (designated 110(1), 110(2), and 110(3))project from the actively cooled electrode 16 (which is of an opensystem type, having outlet apertures 48 for the cooling medium like thatshown in FIGS. 2A/B). Each temperature sensing element 110(1), 110(2),and 110(3) is preferably contained within an isothermal cap 120, aspreviously disclosed and contain a single thermistor 112 (as FIG. 18shows), or multiple spaced-apart thermocouples (in the manner shown inFIGS. 17A/B). The array shown in FIG. 18 allows the physician to obtainand monitor a spatial map of sub-surface temperature conditions aboutthe actively electrode 16. The physician can manually select forablation control purposes the one sensing thermistor (or thermocouple,as the case may be) located in the hottest sub-surface temperatureregion. Alternatively, an automated control mechanism can automaticallycompare temperatures from all thermocouples 110(1), 110(2), and 110(3)and output the hottest sub-surface temperature for temperature controlpurposes. When the multiple-sensor array shown in FIG. 18 is used, theproper orientation of the electrode 16 generally perpendicular to thetissue surface is less critical than when single-sensor embodiments areused.

The embodiment shown in FIG. 20 includes a motor-driven mechanism 162for advancing the stylet 126. In this embodiment, the mechanism 162includes a feedback controller 164 electrically coupled to thetemperature sensing element 110. The feedback controller 164incrementally moves the stylet 126, while taking instantaneousmeasurements of temperature condition at each increment, to seek thesub-surface tissue region where the highest temperature conditionsexist. The controller 164 outputs the sensed highest temperature whileincrementally adjusting the position of the element 110, as necessary,to maintain it in the highest sub-surface temperature region.

Various control processes can be used to command movement of the stylet126 to position the temperature sensing element 110 in the region ofhighest sub-surface tissue temperature. For example, proportionalcontrol principles, proportional integral derivative (PID) controlprinciples, adaptive control, neural network, and fuzzy logic controlprinciples can be used. FIG. 22 shows a representative control process166 that the feedback controller 164 can use.

While incrementally moving the stylet 126, the process 166 inputsinstantaneous tissue temperatures TT(t) sampled by the element 110 at aprescribed time interval Δt. Δt can vary according to the degree ofaccuracy and sensitivity required. For example, Δt can be 5 seconds.

The process 166 derives a temperature difference ΔTT between successivesamples (ΔTT=TT(t)-TT(t-1)). The process 166 employs prescribed courseand fine differential temperature threshold values, respectively E₁ andE₂, to home in on the maximum tissue temperature. The differentialthreshold values can vary, again according to the accuracy andsensitivity required. For example, the course differential thresholdvalue E₁ can be set to 5° C., and the fine differential threshold valueE₂ can be set to 1° C.

As long as ΔTT exceeds the course differential threshold E₁, the process166 commands incremental advancement of the stylet 126, moving theelement 110 deeper into tissue. When ΔTT equals or falls below E₁, theprocess 166 begins to command incremental retraction of the style 126and element 110, while beginning to compare ΔTT to the fine differentialthreshold E₂. The process 166 continues to command incrementalretraction of the stylet 126 as long as ΔTT≦E₁, until ΔTT drops belowE₂, at which time the process 166 commands the stylet 126 to pause forthe set time interval. The process 166 then repeats the above sequence,to seek and maintain the sensor 110 at the depth where the highesttissue temperature exists.

Preferably, the process 166 also sets upper absolute limits foradvancing and retracting the stylet 126 and element 110 within tissue,so that the element 110 remains within a prescribed range of depths toavoid transmural penetration (if too deep) and loss of sub-surfacetissue contact (if not deep enough). Preferably, the speed ofincremental advancement or retraction should be faster than the speed ofthe thermal wave front in the tissue.

The system 148 shown in FIG. 19 is like the system 102 shown in FIG. 9A.As in the system 102, the cooled ablation electrode 16 carries threetemperature sensing elements 104, 106, and 108, for sensing T1, T_(IN),T_(OUT), respectively, as already described. Unlike system 102, in thesystem 148, the cooled ablation electrode 16 carries at least oneadditional temperature sensing element 110 for sensing actual tissuetemperature.

In this arrangement, the master controller 98 receives from thephysician, via the input device 100, a desired tissue temperature valueTT_(SET), a desired electrode temperature T1_(SET), and a P_(MAX). Asearlier disclosed, the set temperature value TT_(SET) represents thedesired hottest sub-surface tissue temperature that the physician wantsto maintain at the ablation site, to thereby control the incidence ofmicro-explosions. TT_(SET) can comprise a fixed value or a set linear ornonlinear curve varying tissue temperature over time.

Likewise, the value T1_(SET) represents a hottest temperature for thethermal mass of the cooled ablation electrode 16, which, as earlierstated, is believed to be between about 10° C. and about 25° C.

The value P_(MAX) is the highest allowed power level, also based uponconsiderations already stated.

The master controller 98 periodically compares the sensed maximum tissuetemperature T_(MAX) to TT_(SET). Based upon this comparison, the mastercontroller 98 derives a demand power output, taking into accountP_(MAX), while cooling to maintain T1_(SET). The demand power outputrepresents the magnitude of the radio frequency power that should besupplied to the electrode 16 to establish and maintain the desiredmaximum tissue temperature TT_(SET).

Alternatively, the master controller 98 could maintain a fixed powerlevel below P_(MAX) and adjust the cooling rate based upon sensedT_(MAX) to achieve TT_(SET). As before described, the master controller98 can control the cooling rate by commanding the temperature controller92 to adjust the temperature of the cooling medium over time, or bycommanding the pump controller 94 to adjust the flow rate of the coolingmedium over time, or by commanding the controllers 92 and 94 in tandemto reach the same result.

The manner in which the controller 98 derives the control commands canvary. For example, it can employ proportional control principles,proportional integral derivative (PID) control principles, adaptivecontrol, neural network, and fuzzy logic control principles. Furtherdetails of these control principle are disclosed in copending U.S.patent application Ser. No. 08/266,023, filed Jun. 27, 1994, andentitled "Tissue Heating and Ablation Systems and Methods UsingTime-Variable Set Point Temperature Curves for Monitoring and Control."

In a preferred implementation, the controller 98 sets a value for Åbased upon the magnitude of the current demand power value, as set bythe sensed tissue temperature condition T_(MAX). The controller thencontrols the cooling rate to achieve the set value for Å. In this way,the controller maximizes the benefits of cooling the electrode at thedemand power value.

The illustrated and preferred embodiments envision the use ofmicro-processor controlled components using digital processing toanalyze information and generate feedback signals. It should beappreciated that other logic control circuits using micro-switches,AND/OR gates, invertors, and the like are equivalent to themicro-processor controlled components and techniques shown in thepreferred embodiments.

Various features of the invention are set forth in the following claims.

We claim:
 1. A system for ablating body tissue comprisingan ablation electrode having a thermal mass, the ablation electrode having a tissue contact region for contacting a tissue surface, the ablation electrode adapted to be connected to a source of radio frequency energy to conduct radio frequency energy for transmission by the ablation electrode into the tissue surface contacting the tissue contact region, the ablation electrode including an interior well in the tissue contact region, a tissue temperature sensing assembly carried within the interior well comprising a temperature sensor, a cap enclosing the temperature sensor, the cap being made of a thermal conductive material, which is in thermal conductive contact with the temperature sensor, the thermal conductive material having a thermal conductivity that is at least 1.0 W/m K, and an insulating barrier made of thermal insulating material in the interior well to substantially thermally isolate the cap from the thermal mass of the ablation electrode, the cap having a distal end, a mechanism attached to the cap to selectively move the cap in the interior well to locate the distal end of the cap beyond the tissue contact region of the ablation electrode and into thermal conductive contact with tissue at different distances beneath the tissue surface, the thermal conductive material of the cap reaching thermal equilibrium with tissue temperature conditions without dissipation by the termal mass of the ablation electrode, and a controller coupled to the mechanism and to the temperature sensor to control movement of the cap in the interior well based, at least in part, upon tissue temperature conditions sensed by the temperature sensor within the cap.
 2. A system according to claim 1wherein the controller is configured to control the mechanism to locate the distal end of the cap at a distance beneath the tissue surface, where the temperature sensor senses a maximum tissue temperature condition.
 3. A system according to claim 1 or 2wherein the temperature sensor element comprises a thermistor.
 4. A system according to claim 1 or 2wherein the temperature sensor element comprises a thermocouple.
 5. A system according to claim 1 or 2wherein the mechanism is configured to rotate the cap in the interior well to thereby move the distal end of the cap beyond the tissue contact region of the ablation electrode.
 6. A system according to claim 1 or 2wherein the mechanism is configured to move the distal end of the cap beyond the tissue contact region of the ablation electrode without rotation of the cap in the interior well.
 7. A system according to claim 1 or 2wherein the distal end of the cap is blunt and is configured to project into tissue below the tissue surface during movement of the cap.
 8. A system according to claim 1 or 2wherein the distal end of the cap includes a sharpened point, which is configured to penetrates tissue below the tissue surface during movement of the cap.
 9. A system for ablating body tissue comprisinga generator for supplying radio frequency ablation energy, an electrode having a thermal mass, the ablation electrode having a tissue contact region for contacting a tissue surface, the electrode and the generator being coupled to conduct radio frequency ablation energy for transmission by the electrode into the tissue surface contacting the tissue contact region, the ablation electrode including an interior well in the tissue contact region, a tissue temperature sensing assembly carried within the interior well comprising a temperature sensor, a cap enclosing the temperature sensor, the cap being made of a thermal conductive material, which is in thermal conductive contact with the temperature sensor, the thermal conductive material having a thermal conductivity that is at least 1.0 W/m K, and an insulating barrier made of thermal insulating material in the interior well to substantially thermally isolate the cap from the thermal mass of the ablation electrode, the cap having a distal end, a mechanism attached to the cap to selectively move the cap in the interior well to locate the distal end of the cap beyond the tissue contact region of the electrode and into thermal conductive contact with tissue at different distances beneath the tissue surface, the thermal conductive material of the cap reaching thermal equilibrium with tissue temperature conditions without dissipation by the termal mass of the ablation electrode, a first controller coupled to the mechanism and to the temperature sensor to control movement of the cap in the interior well based, at least in part, upon tissue temperature conditions sensed by the temperature sensor within the cap, and a second controller coupled to the temperature sensor and to the generator to control the supply of radio frequency ablation energy based, at least in part, upon temperature conditions sensed by the temperature sensor within the cap.
 10. A system for ablating body tissue comprisinga generator for supplying radio frequency ablation energy, an electrode having a thermal mass, the ablation electrode having a tissue contact region for contacting a tissue surface, the electrode and the generator being coupled to conduct radio frequency ablation energy for transmission by the electrode into the tissue surface contacting the tissue contact region, the ablation electrode including an interior well in the tissue contact region, an element to cool the electrode, a tissue temperature sensing assembly carried within the interior well comprising a temperature sensor, a cap enclosing the temperature sensor, the cap being made of a thermal conductive material, which is in thermal conductive contact with the temperature sensor, the thermal conductive material having a thermal conductivity that is at least 1.0 W/m K, and an insulating barrier made of thermal insulating material in the interior well to substantially thermally isolate the cap from the thermal mass of the ablation electrode, the cap having a distal end, a mechanism attached to the cap to selectively move the cap in the interior well to locate the distal end of the cap beyond the tissue contact region of the electrode and into thermal conductive contact with tissue at different distances beneath the tissue surface, the thermal conductive material of the cap reaching thermal equilibrium with tissue temperature conditions without dissipation by the termal mass of the ablation electrode, a first controller coupled to the mechanism and to the temperature sensor to control movement of the cap in the interior well based, at least in part, upon tissue temperature conditions sensed by the temperature sensor within the cap, and a second controller coupled to the temperature sensor and to the generator to control the supply of radio frequency ablation energy based, at least in part, upon temperature conditions sensed by the temperature sensor within the cap.
 11. A system for ablating body tissue comprisinga generator for supplying radio frequency ablation energy, an electrode having a thermal mass, the ablation electrode having a tissue contact region for contacting a tissue surface, the electrode and the generator being coupled to conduct radio frequency ablation energy for transmission by the electrode into the tissue surface contacting the tissue contact region, the ablation electrode including an interior well in the tissue contact region, an element to cool the electrode, a tissue temperature sensing assembly carried within the interior well comprising a temperature sensor, a cap enclosing the temperature sensor, the cap being made of a thermal conductive material, which is in thermal conductive contact with the temperature sensor, the thermal conductive material having a thermal conductivity that is at least 1.0 W/m K, and an insulating barrier made of thermal insulating material in the interior well to substantially thermally isolate the cap from the thermal mass of the ablation electrode, the cap having a distal end, a mechanism attached to the cap to selectively move the cap in the interior well to locate the distal end of the cap beyond the tissue contact region of the electrode and into thermal conductive contact with tissue at different distances beneath the tissue surface, the thermal conductive material of the cap reaching thermal equilibrium with tissue temperature conditions without dissipation by the termal mass of the ablation electrode, a first controller coupled to the mechanism and to the temperature sensor to control movement of the cap in the interior well based, at least in part, upon tissue temperature conditions sensed by the temperature sensor within the cap, and a second controller coupled to the temperature sensor and to the cooling element to control the cooling of the electrode based, at least in part, upon temperature conditions sensed by the temperature sensor within the cap.
 12. A system according to claim 9 or 10 or 11wherein the first controller is configured to control the mechanism to locate the distal end of the cap at a distance beneath the tissue surface, where the temperature sensor senses a maximum tissue temperature condition.
 13. A system according to claim 9 or 10 or 11wherein the temperature sensor comprises a thermistor.
 14. A system according to claim 9 or 10 or 11wherein the temperature sensor comprises a thermocouple.
 15. A system according to claim 9 or 10 or 11wherein the mechanism is configured to rotate the cap in the interior well to thereby move the distal end of the cap beyond the tissue contact region of the ablation electrode.
 16. A system according to claim 9 or 10 or 11wherein the mechanism is configured to move the distal end of the cap beyond the tissue contact region of the ablation electrode without rotation of the cap in the interior well.
 17. A system according to claim 9 or 10 or 11wherein the distal end of the cap is blunt and projects into tissue below the tissue surface during movement of the cap.
 18. A system according to claim 9 or 10 or 11wherein the distal end of the cap includes a sharpened point that penetrates tissue below the tissue surface during movement of the cap. 